The structure and components of soft tissues are discussed in nearly any textbook on human physiology (e.g., Guyton and Hall, Textbook of Medical Physiology, 9th edition (1996) at page 186). Very briefly, the cells in most types of “soft tissue” (excluding bones, teeth, fingernails, etc.) are held together by a matrix (i.e., a three-dimensional network) of two types of fibers.
The larger and stronger type of fiber is collagen, a well-known fibrous protein that provides most of the tensile strength of tissue.
The other major type of fiber that helps hold together soft tissue consists mainly of “proteoglycan filaments”. These filaments contain a small quantity of protein and a much larger amount (roughly 98%) of hyaluronic acid, a natural polymer with alternating saccharide rings of glucosamine and glucuronate. Unlike collagen fibers, which are thick and provide high levels of tensile strength, proteoglycan filaments are extremely thin, and cannot be seen under light microscopes. They cause the watery extra-cellular fluid in soft tissue to form a gel-like material called “tissue gel”. This gel contains water, the proteoglycan filaments, and any other extra-cellular molecules that are suspended in the watery solution.
Roughly ⅙ of the volume of a person's body is made up of tissue gel, and it is essential to proper functioning of any type of soft tissue; among other things, it helps oxygen and nutrients reach cells, it aids in the removal of waste metabolites from tissue, and it helps tissue remain flexible and supple. Because proteoglycan filaments are so thin, molecules dissolved in tissue gel can permeate through the gel material with very little impedance; experiments have indicated that dye molecules can diffuse through tissue gel at rates of about 95 to 99 percent of their diffusion rates in water or saline.
Because nearly any type of soft tissue, in its normal and natural state, can be regarded as a type of hydrogel, many efforts have been made to create and use synthetic polymeric hydrogel materials as tissue implants. Most of these polymers are created by using non-parallel strands of long organic polymeric molecules (usually with chemical structures that are easier to work with and manipulate than glucosamine and glucuronate). Such molecules, to be suitable for use in a hydrogel, must be very hydrophilic (i.e., they must be able to attract and hold large quantities of water). This is most frequently accomplished by polymerizing precursor molecules that will provide large numbers of hydroxy groups (or other hydrophilic groups), on relatively short “side chains” or “side groups” that are bonded in a regular spaced manner to the long “backbone” strands of the final polymer.
An example of a synthetic hydrogel of this nature is PHEMA (an acronym for poly-hydroxy-ethyl-meth-acrylate), which is used to make soft contact lenses, drug-releasing hydrogels, and similar articles. In contact lenses made of PHEMA, the polymer does not actually bend light. Instead, the water that dwells inside the PHEMA polymer when the lens is hydrated does that job. The hydrophilic PHEMA polymer merely holds water molecules together, in the shape of a contact lens. If a polymer such as PHEMA is dehydrated, it typically becomes brittle; as long as it remains filled with water, it stays soft and flexible. However, like most synthetic hydrogels, PHEMA does not have sufficient strength and durability to last for years (or decades) as a permanent surgical implant.
PHEMA is not the only synthetic polymer used to create biocompatible hydrogels; other polymers that can swell and soften when saturated with water include various hydrophilic polyurethane compositions (e.g., Gorman et al 1998 and U.S. Pat. No. 4,424,305, Gould et al 1984), poly(vinyl alcohol) compositions (e.g., Wang et al 1999), polyacrylonitrile compositions (e.g., U.S. Pat. No. 4,420,589 (Stoy 1983) and U.S. Pat. No. 4,493,618 (Stoy et al 1990)), and other compounds known to those skilled in this field of art.
The flexible, pliable, gel-like nature of a water-saturated synthetic hydrogel arises mainly from crosslinking attachments between non-parallel fibers in the gel. Depending on the specific polymeric structure that has been chosen, the crosslinking attachments between the long “backbone” chains in a hydrogel polymer can be formed by covalent bonding, by hydrogen bonding or other ionic attraction, or by the entanglement of chains that have relatively long and/or “grabby” side-chains or groups. Regardless of which type of bonding or entangling method is used to hold the polymeric chains together in a hydrogel, the “coupling” points between the chains can usually be flexed, rotated, and stretched.
In addition, it should be recognized that the backbone chains in hydrogel polymers are not straight. Instead, because of certain aspects of interatomic bonds and the “tetrahedral” arrangement of the electron pairs in the valence shells of carbon atoms, the long backbone chain of a polymeric molecule will be kinked, and can be stretched somewhat, in an elastic and springy manner, without breaking the bonds.
In a typical hydrogel, the fibers take up only a small portion of the volume (usually less than about 10%, and many hydrogels contain less than 2% fiber volume). The large majority (usually at least 90 to 95%) of the volume is made up of “interstitial” spaces (i.e., the unoccupied spaces that are nestled within the three-dimensional network of fibers, and that become filled with water, when a gel is hydrated). Accordingly, since the “coupling” point between any two polymeric backbone chains can be rotated and flexed, since polymeric backbone molecules can be stretched somewhat without breaking them, and since the water molecules contained in a hydrogel effectively provide a free-flowing lubricant that will not impede relative motion between adjacent polymeric chains, a supple and resilient gel structure is created, when a synthetic hydrogel polymer is hydrated.
Various methods are known for creating conventional polymeric hydrogels. A number of such methods involve mixing together and reacting precursor materials (monomers, mixtures of polymer strands and cross-linking agents, etc.) while they are suspended in water or other solvent. The process of forming cross-linked polymeric strands, while they are suspended in a solvent, can be used to give a desired density and three-dimensional structure to the resulting polymerized matrix. The resulting material is then frozen, to preserve the desired three-dimensional structure of the matrix, and the ice (or other frozen solvent) is then removed, using a process called sublimation or lyophilization. This process uses a combination of freezing temperature and intense vacuum, to cause the frozen solvent to vaporize without passing through a liquid stage, since passage through a liquid stage might alter and damage the three-dimensional molecular structure that will give the gel its final desired shape, cohesion, and other properties.
After the solvent has been removed, any final steps (such as a final crosslinking reaction, rinsing or washing steps to remove any unreacted reagents, etc.) are carried out. The polymer is then gradually warmed up to room temperature, and it is subsequently saturated with water, to form a completed hydrogel.
These and other methods for creating synthetic polymeric hydrogels that are biocompatible and intended for surgical implantation are described in numerous patents, including U.S. Pat. No. 3,822,238 (Blair et al 1974), U.S. Pat. No. 4,107,121 (Stoy 1978), U.S. Pat. No. 4,192,827 (Mueller et al 1980), U.S. Pat. No. 4,424,305 (Gould et al 1984), U.S. Pat. No. 4,427,808 (Stol et al 1984), and U.S. Pat. No. 4,563,490 (Stol et al 1986). In addition, various methods of forming hydrogel coatings on the surfaces of other (“substrate”) materials are also described in various patents, such as U.S. Pat. No. 4,921,497 (Sulc et al 1990) and U.S. Pat. No. 5,688,855 (Stoy et al 1997).
There also have been efforts to reinforce some types of hydrogels with an interpenetrating network (IPN) of fibers, to enhance and strengthen a hydrogel's mechanical properties. These reinforcing fibers are usually either chopped into fragments, or longitudinally aligned with the fibers within the hydrogel. A number of these efforts to develop “composite hydrogels” apparently have focused on attempts to create synthetic pericardial tissue (i.e., the membrane that surrounds the heart); see, e.g., Blue et al 1991 and Walker et al 1991. Articles which describe these and other efforts to develop “composite” hydrogels that can be surgically implanted in humans are discussed in two review articles, Corkhill et al 1989 and Ambrosio et al 1998.
As used herein, all references to “implants” or “implantation” (and all terms such as surgery, surgical, operation, etc.) refer to surgical or arthroscopic implantation of a reinforced hydrogel device, as disclosed herein, into a mammalian body or limb, such as in a human patient. Arthroscopic methods are regarded herein as a subset of surgical methods, and any reference to surgery, surgical, etc., includes arthroscopic methods and devices. The term “minimally invasive” is used occasionally, but it is imprecise; one should assume that any surgical operation will be done in a manner that is minimally invasive, in view of the status and needs of the patient, and the goals of the surgery.
Fibro-Cartilage Implants in Spinal Disc Repair
One subset of efforts to develop composite implants with embedded fibers merits attention, but only to point out that it is not relevant to this invention.
A number of researchers have attempted to develop improved fiber-reinforced hydrogel materials, for use in making artificial “intervertebral discs”, for repairing or replacing damaged spinal discs. Those efforts are not relevant herein, and the reasons for that exclusion should be recognized, for a better understanding of this invention.
Briefly, the hard vertebral bones of the spine are separated from each other by somewhat flexible discs. Those discs are made of a type of highly fibrous cartilage that is called “fibro-cartilage” in most medical texts (although some books, articles, and patents refer to hyaline cartilage on vertebral discs). The fibro-cartilage material in spinal discs is very different from the type of cartilage that provides the smooth sliding surfaces that cover the sliding ends of bones in joints such as knees, hips, shoulders, etc.
The fibro-cartilage material found in spinal discs must be able to flex, compress, and rotate somewhat, when a person bends over or twists his body. However, spinal discs do not have smooth, slidable, lubricated, articulating surfaces, because they emphatically must not allow any sliding motion to occur, between two adjacent vertebral bones. Instead, one of their primary functions is to totally prevent and prohibit any sliding motion between adjacent vertebral bones. Any such sliding motion, if allowed to occur in the spinal column, would translate directly into shearing motion, and shearing stresses. If allowed to occur in the spine, that type of relative motion between adjacent vertebral bones would pose a grave risk of pinching, injuring, and even shearing (i.e., transversely cutting) the spinal cord.
Therefore, the fibro-cartilage material in spinal discs evolved in ways that prevent sliding motions. Instead of having smooth and slippery surfaces that actively promote smooth sliding motions between adjacent bones, a spinal disc has fibers that extend outwardly from both sides of the disc. Those fibers extend roughly a quarter of an inch into vertebral bones, in a manner that forms transition zones, where fibers from a spinal disc are interlaced with crystals of hydroxyapatite, the calcium-phosphate mineral structure that forms hard bones. Those fiber/mineral transition zones are crucial in reinforcing the attachment of a spinal disc to its adjacent vertebral bones, and in preventing any form of sliding motion that could injure the spinal cord.
Hyaline (or “articulating”) cartilage, of the type that covers a bone surface in a joint such as knee or hip, has a totally different structure, because it serves a totally different purpose. Unlike the cartilage in spinal discs, which evolved in ways that absolutely prevent and prohibit any sliding motion, hyaline cartilage in articulating joints absolutely must have at least one smooth and slippery surface, in order to enable the type of sliding motion that occurs in joints such as knees, shoulders, and hips.
Meniscal cartilage (in knee joints) and labral cartilage (in hip and shoulder joints) is in a curious intermediate category. As with hyaline cartilage, meniscal and labral cartilage must provide a smooth and slippery surface, which is specifically designed to promote sliding motion and articulation of those joints. This crucial factor renders meniscal or labral cartilage completely and totally different from spinal discs.
However, it also should be noted that meniscal and labral cartilage have fibrous internal structures, and are usually classified as fibro-cartilage. Unlike hyaline cartilage, they do not form a relatively thin layer that is anchored to a hard bone; instead, meniscal and labral cartilage segments are formed in relatively long semi-circular arcs, which are anchored to bone only through ligaments that extend out of both tips of the arc. Therefore, the internal structure of meniscal or labral cartilage requires high levels of internal fibrous reinforcement. This led to the evolution of fibro-cartilage internal structures. However, as with hyaline cartilage, this fibrous reinforcing structure is covered with smooth and slippery surfaces; accordingly, it is totally different from the fibrous interfaces that must totally prevent any slippage between vertebral bones and spinal discs, in a mammalian spine.
Accordingly, prior art efforts to develop fiber-reinforced cartilage replacements, for use in repairing or replacing spinal discs (as described in various patents such as U.S. Pat. No. 4,911,718, Lee et al 1990, and U.S. Pat. No. 5,171,281, Parsons et al 1992), are worth noting, but only in passing. This current invention has a completely different goal, and a completely different set of physical and performance constraints. Instead of creating spinal disc replacements that will carefully prevent any sliding motion, to avoid shearing stresses on a spinal cord, the goal of this invention is to provide cartilage segments that will provide the exact opposite function: they must have extremely smooth and “lubricious” surfaces, to promote smooth and lubricated sliding motions (also called “articulating” motion) between adjacent bones, in joints such as knees, hips, and shoulders.
The challenge of providing a smooth and lubricated surface, in a surgical implant that will replace a segment of articulating cartilage, can be accomplished most effectively through the use of hydrogel materials. However, hydrogel materials are notoriously weak and easy to damage, due to the fact that they contain a large majority (by volume) of water molecules, and only a small quantity of polymer or protein molecules that will provide the three-dimensional matrix that holds the water molecules inside the hydrogel.
Accordingly, this invention discloses special types of reinforced hydrogel implants having smooth and lubricious surfaces, for replacing cartilage in articulating joints.
Resorbable Versus Non-Resorbable Implants
It must also be recognized that the reinforced hydrogel implants disclosed herein fall into the “non-resorbable” category of cartilage implants. These implants are designed to entirely replace a segment of damaged or diseased cartilage, on a permanent basis. As such, they are designed to remain inside the joint of the patient for multiple years (preferably for the entire remaining life of the patient, if possible).
This trait is important, because it excludes implants that are designed to hold and protect transplanted cells that will generate new cartilage. Numerous researchers (including the Applicant herein) are working hard to try to develop “resorbable” implants, which involve matrices made of collagen fibers, polyhydroxyalkanoate polymers, or other resorbable materials. These matrices can shelter and protect certain types of specialized transplanted cells (such as chondrocyte cells, and mesenchymal stem cells) that will generate new biological cartilage over a span of weeks or months. These types of resorbable will gradually be dissolved by bodily fluids, over a span of months, in a manner similar to the gradual dissolving of collagen fibers, which are constantly being recycled and replaced in any mammalian body by new collagen fibers that are gradually secreted by certain types of cells in connective tissue.
Surgical implants that are made of collagen fibers or resorbable polymers, and that are used to nurture and protect transplanted cells that generate new biological cartilage, can also be referred to as “biological” implants, to distinguish them from non-resorbable implants made of synthetic materials. They are discussed in more detail in numerous articles and patents, including U.S. Pat. No. 6,530,956 (by Mansmann, the same Applicant herein). However, those types of resorbable implants are not discussed in any more detail herein, because they are not relevant to this current invention. Instead, this invention relates solely to non-resorbable implants that are designed to be permanent, and that will not be absorbed by the body, even after numerous years.
The types of cartilage replacement implants that are disclosed herein are designed for two distinct types of uses. One category will be anchored to hard bone structures, and will be referred to herein, for convenience, as “condylar” implants. The other category will be anchored to soft tissue rather than hard bone structures, and will be referred to herein, for convenience, as “meniscal” implants.
These two different types of implants will both require fiber-reinforced hydrogels, with smooth articulating surfaces. However, since they will require different types of anchoring structures, they are discussed under separate subheadings below.